Cardiac assistance systems having bi-directional pumping elements

ABSTRACT

A unified, non-blood contacting, implantable heart assist system surrounds the natural heart and provides circumferential contraction in synchrony with the heart&#39;s natural contractions. The pumping unit is composed of adjacent tube pairs arranged along a bias with respect to the axis of the heart and bound in a non-distensible sheath forming a heart wrap. The tube pairs are tapered at both ends such that when they are juxtaposed and deflated they approximately follow the surface of the diastolic myocardium. Inflation of the tube pairs causes the wrap to follow the motion of the myocardial surface during systole. A muscle-driven or electromagnetically powered energy converter inflates the tubes using hydraulic fluid pressure. An implanted electronic controller detects electrical activity in the natural heart, synchronizes pumping activity with this signal, and measures and diagnoses system as well as physiological operating parameters for automated operation. A transcutaneous energy transmission and telemetry subsystem allows the Unified System to be controlled and powered externally.

GOVERNMENT FUNDING

The work described herein was supported, in part, by U.S. GovernmentGrant Number NIH-NOI-HV-58154 awarded by the National Institutes ofHealth. The Government, therefore, may have certain rights in thisinvention.

BACKGROUND OF THE INVENTION

The present invention relates generally to cardiac assist and/orresuscitation systems for restoration or augmentation of natural bloodflow and, more particularly, to implantable systems and associatedmethods for assisting the natural contractions of the heart.

Following a heart attack or as a result of other cardiac disease states,the heart's ability to pump blood can be seriously impaired.Conventional cardiac assist systems employ a variety of pumpingapproaches for assisting a failing natural heart. Generally, there aretwo categories of cardiac assist systems: those which contact blood,referred to herein as blood-contacting cardiac assist systems; and thosewhich do not, referred to herein as non-blood-contacting cardiac assistsystems.

A primary drawback of blood-contacting cardiac assist systems is theassociated risk of thromboembolism. Although significant efforts havebeen made to reduce or eliminate this problem, the continued risk ofthrombosis has restricted blood-contacting cardiac support systems totemporary or short-term applications. In addition to the risk ofthrombosis, blood-contacting cardiac assist devices typically alsoexperience calcification. The degree of calcification increases withtime, again making these devices undesirable for long term applications.

Non-blood-contacting cardiac support systems significantly reduce therisk of thromboembolism and calcification. One conventional approach hasbeen to directly apply forces to the heart so as to facilitate pumping.For example, U.S. Pat. No. 4,304,225 to Freeman discloses anon-contacting cardiac assist system designed to compress all or part ofthe heart by alternately tightening and releasing a circumferentialcompression band. Another conventional device, described in U.S. Pat.No. 4,583,523 to Kleinke et al., is an articulated mechanical device forapplying an encircling force to the aorta. European Publication No.0583012 A1 to Heilman et al. teaches the application of a similar deviceto the heart. Still other conventional systems, such as U.S. Pat. No.4,411,268 to Cox and U.S. Pat. No. 4,813,952 to Khalafalla disclose anapproach of encircling the heart with the latissimus dorsi muscle toachieve a desired compression of the heart.

Another class of non-blood-contacting cardiac assist devices useshydraulic or gas pressure to displace an equivalent volume of blood inthe heart through pressure applied to the outer surface of the heart,the epicardium. One conventional approach has been to use a housing ofrigid construction for enveloping, at least partially, the ventricularregion of the myocardium. The inner surface of the housing typically hasa distensible elastic membrane adjacent to the myocardial wall. Pumpingfluids are fed to the chambers defined by the housing and the membraneto apply pressure on the myocardial wall. In some instances the outerportion of the housing is formed of a flexible, non-distensible, memberwith an elastic distensible inner membrane. In general theseconventional approaches utilize one or more compartments, eachcharacterized by an elastic inner wall and an inelastic outer wall.Filling the compartments compresses the myocardium of the ventricle toaid pumping. When pumping is facilitated in this manner, a volume ofinflating fluid or gas is required to displace an equal volume of blood.Cardiac assist devices of this general class are described in U.S. Pat.Nos. 2,826,193; 3,371,662; 3,455,298, 3,587,567; 3,613,672; 4,048,990;4,192,293; 4,506,658; 4,536,893; 4,690,134; 4,731,076; 5,119,804;5,131,905; 5,169,381; and 5,273,518.

Other approaches have employed a concave, gel-filled compression padactivated by a plate on its outer surface (U.S. Pat. Nos. 4,925,443;5,098,369; 5,348,528); a cardiac assist envelope designed for minimallyinvasive implantation (U.S. Pat. No. 5,256,132); or a cardiac assistdevice having a fluid filled jacket encasing at least the heartventricles to provide a compliant, completely passive support (U.S. Pat.No. 4,957,477).

A drawback to these cardiac assist devices is that they typically pumpblood by displacing the blood with an equal inflation volume of ahydraulic fluid. As a result of this limitation, such systems requirelarge reservoirs of the hydraulic fluid and/or complex pumpingprotocols.

To overcome this and other drawbacks, cardiac assist devices have beendevised which displace blood with an inflation volume smaller than thedisplaced blood volume. Such cardiac assist devices typically producehigher pumping capacities through the injection of a relatively smallerquantity of fluid or gas under high pressure. Generally, these devicesutilize a chamber or wrap having a number of inflatable segments.

For example, commonly owned U.S. Pat. No. 5,713,954 to Rosenberg et al.describes a non-blood-contacting cardiac assist device having tubes thatcontract a circumference of the heart when inflated. In one embodiment,the Rosenberg device is constructed of vertically-oriented, cylindrical(tube-shaped) inflation chambers arranged to form a ring and surroundedby a nondistensible sheath to form an artificial myocardium or heartwrap. The administration of a fluid under pressure causes the tubes ofthese conventional devices to have an expanded cross section, which isgenerally circular. When the fluid is withdrawn, the tubes flattenperpendicular to the direction of force generated by the pressure in theheart. When the tubes are deflated, the circumference of the pumpingchamber is equivalent to the value of the number of tubes in the wrapmultiplied by one half the circumference of one of the constituenttubes. When the tubes are fully inflated, the circumference of thepumping chamber is equivalent to the product of the number of tubes andthe inflated diameter of one of the constituent tubes. When the wrapcircumference is minimized there is no dilation of the tubecircumference.

The resulting contraction of the circumference of the heart wrap ismaximally 36%. This limit is due to the geometry of the device and isindependent of the radius of the tubes chosen. Therefore, the volume ofeach tubes can be made small while maintaining a constant ejectionvolume. However, the work done is, in all cases, the same. The result isthat smaller tubes require a higher pressure to attain a circular crosssection. In general, for constant work preformed, the inflation pressureis inversely proportional to the inflation volume.

U.S. Pat. No. 3,464,322 to Pequigot also discloses an artificial bloodpumping chamber that has walls which are formed from an arrangement ofinflatable tubes. A drawback of the Pequigot device is that theinflation chamber tubes are free to dilate when inflated. The Rosenbergdevice overcomes the drawbacks of the Pequigot device since thecircumference of the inflation chambers of the Rosenberg device cannotexceed the dimensions of the fabric pockets in which they are imbedded.Consequently, the pumping action resulting from a contraction of theRosenberg heart wrap is not defeated by dilation of the radii of theinflation chambers. Therefore, the Rosenberg device is more likely toreach the theoretical maximum contraction of 36%. However, like thePequigot device, the Rosenberg device cannot exceed this limitingmaximum contraction ratio.

One drawback with the above blood pumping devices is that the actualextent of contraction, expressed as a percentage of the circumference ofthe deflated pumping chamber, is dependent upon the amount ofnon-contracting space between the tubes. However, in practice, it isvery difficult to inflate a sheet of tubes, joined only at a tangent,without inducing high stress in the tubes or in an encircling sheath.Maintenance of the tubes in close proximity at high pressurenecessitates some non-contracting space between the tubes. Furthermore,since the pumping chamber is meant to fit snugly to the heart,allocation must be made for fitting the pumping chamber to the heart insitu. Consequently, the tubes must be spaced apart for this purpose. Asa result, the ejection volumes produced by the heart as a result of thespacing apart of the tubes in these conventional devices aresignificantly less than desired. This drawback occurs even if theinflated portion of the pumping chamber's circumference were to produceits theoretical maximum of 36% diametric contraction.

What is needed, therefore, is a non-blood contacting ventricular assistdevice that generates a contraction, which exceeds the theoretical limitof conventional contractile balloon wraps, and hence, generates agreater maximum stroke volume. The device should not encumber thenatural function of the heart and, in the event of failure, the deviceshould not interfere with the natural pumping action of the heart. Thepossibility of further injury to the heart and adjacent vessels shouldalso be minimized by providing gentle and physiologically correctpumping action. The device should not damage adjacent tissue ortraumatize adjacent organs by compression or excessive localizedtemperatures. The ventricular assist device should also be configurableto assist the left, right or both ventricles.

In addition to cardiac assist devices which actively assist the heart inpumping blood (so-called “active devices”), the present invention alsopertains to another type of heart assist device known as a “passiveconstraint” or simply a “passive” device. Passive devices serve toprevent cardiac expansion beyond a predetermined volumetric limit inpatients suffering from cardiac dilation, hypertrophy and relatedconditions. In the absence of such constraint, the weakened heart musclewill lose its ability to pump blood and, in many cases, result in damageto the patient's heart valves. In passive devices, the goal is not toaugment or replace the natural heart's pumping action but rather toassist the heart by applying a constraining force during the heart'sexpansion (diastolic) phase.

Ideally, a passive device wrapped around the heart should mimic thenatural resistance of the heart muscle itself to over-expansion. Ahealthy natural heart will exhibit a characteristic relationship betweenventricular pressure and volume, such that small amounts of pressurewill initially result in a desired expansion of the ventricular volume.During activity or exercise, the ventricles must also response to higherpressures to accommodate a greater volumetric expansion and, thereby,permit increased ventricular output. However, to achieve a desiredventricular output, especially during normal activity or exercise, themaximum ventricular end-diastole volume must still be constrained orelse the ventricle, during contraction, will be unable to eject thenecessary quantity of blood.

Unfortunately, conventional passive devices exhibit constraining forcesthat typically are not well matched to the natural physiology. Even whensuch passive wraps are constructed from elastic materials that respondto increases in pressure in accordance with Young's law, the performanceof such devices degrades over time, largely due to the in-growth ofepicardial and/or interstitial cells within and around the device. Thisin-growth prevents the elastic elements of the device from stretching.

Thus, what is also needed is a passive cardiac device that can bettermimic the natural heart's response to increases in diastolic pressureand, in particular, devices that can continue to function and respond tosuch pressures despite in-growth of cells over time.

SUMMARY OF THE INVENTION

Methods and apparatus are disclosed for providing assistance to theventricles of a natural heart. In one aspect of the invention,ventricular assist devices, capable of encircling at least a portion ofthe heart, are disclosed having multiple layers of inflatable elements.Such multi-layer devices induce contractions that overcome the limits ofconventional contractile balloon wraps, and hence, generate a greatermaximum stroke volume. Pumping modules incorporating multiple layers ofinflatable elements are disclosed to wrap around, or attach to, one orboth ventricles of a natural heart, or to any-other blood containingstructure that enables natural circulation. The invention alsoencompasses a unified system that integrates the pumping modules of thepresent invention with other major components required for mechanicalheart massage into one system that is completely implantable into thehuman thorax.

Thus, a unified system according to the present invention can becomposed of a highly efficient pumping module (described in more detailbelow), one or more reservoirs of fluid and a control module. Thecontrol module can include an internal electronic controller forgenerating a suitably shaped pressure wave to be synchronized with thenatural contraction of the heart, pumps, valves and/or regulators fordelivering a pressurized fluid to the pumping module and a suitablepower supply. In one embodiment, the power supply can include atranscutaneous energy-receiving device, and/or an implantable batteryfor storage of the received energy. The control module can furtherinclude a data transceiver.

The pumping module can also include conduits for distributing thepressurized fluid to the various inflatable elements of the pumping unitand attachment elements for attaching the pumping unit to the heart. Theattachment mechanism can include direct attachment elements or tetheringdevices intended to prevent the wrap from slipping off the heart.

In accordance with another aspect of the invention, an easily attachable(and in at least some instances, a readily detachable) pumping unit isdisclosed that is constructed of thin, collapsible, non-distensible,biocompatible material, which encases a multi-layer arrangement ofinflatable elements. The inflatable elements can also be bound by asheath that holds them in a defined geometry. For example, in onearrangement, the encircling sheath can bind sets of two or moreinflatable elements in individual pockets, such that when the sheath iswrapped about a heart, the inflatable elements of each pocket will bestacked or juxtaposed along a radial line. The sheath also serves tojoin the sets of inflatable elements to each other along a lineperpendicular to said radial line. This second dimension thus forms acircumferential restriction when placed around the heart.

The inflation elements can be tapered at one or both ends, and theresulting wrap curved in a plane, so that when joined end-to-end to forma continuous band, the wrap describes approximately the surface of aparaboloid of revolution. In this configuration, the surface of wrapthat faces the epicardium of the heart presents a plurality of pocketseach of which contain multiple layers of inflatable elements. Theinflatable elements can be filled at time of implantation to conform thewrap to the heart. In one embodiment, particularly useful in activedevices that are intended to assist the natural heart's pumping action,the inflatable elements are filled with a flexible, deformable substancethat substantially maintains its volume when compressed.

Unified systems according to the invention can further include one ormore electrodes for implantation on the heart or at a suitable adjacentsite (e.g., on the heart assist device), to sense the heart's electricalsignals and synchronize pump activation with the heart's cycle. Forexample, the sensor can detect and/or monitor well-known EKG componentssuch as the p-wave, or the q-r-s-wave (indicating the beginning ofsystole) and/or the t-wave (at the end of systole). The signals fromsuch sensor electrodes are then used by an electronic controller tosynchronize the release of actuating fluid to the pumping unit and,subsequently, to synchronize the evacuation of fluid from the pumpingunit.

The unified system can further include an energy converter (e.g., apump) and at least one plenum for storage of a pressurized volume offluid of sufficient size to provide a flow at nearly constant pressureduring systole and to provide a flow away from the heart assist deviceat nearly constant vacuum (i.e., at a pressure less than ambient) duringdiastole.

Fluid control can be accomplished in at least two different ways. In oneembodiment, a single plenum is used to store the inflation fluid and abi-directional constant pressure pump is used to both inflate andevacuate the heart assist device. The pump can be a electromechanicalenergy converter or a device that induced the patient's own skeletalmuscles to power a pump, or it can be a hybrid of both. Alternativelytwo uni-directional pumps can be used in tandem to fill and empty theinflatable elements of the heart assist device.

In a second embodiment, two plenums can be employed. One plenum isprovided for storage of evacuated fluid and is preferably maintained ata sufficient state of evacuation so as to provide evacuating flow at anearly constant pressure during the evacuation interval. A second plenumis provided for filling the inflatable elements of the heart assistdevice and is preferably maintained at a sufficient state ofpressurization so as to fill the heart-pumping unit at a nearly constantpressure during systole. The unified system can further include amechanical device or energy converter for continuously pumping fluidfrom the evacuated plenum to the pressurized plenum. The system canfurther include a controller or regulators to maintain the plenums attheir respective pressurized states. In one embodiment, the energyconverter, controller and plenums are contained in a single housing, theback of which has a convex surface curvature compatible with theinternal abdominal cavity.

The system can further comprise a mechanical device or energy converterfor continuously pumping fluid from the evacuated plenum to the lowpressurize plenum and maintaining their respective pressurized states,and a second energy converter for continuously pumping fluid from theevacuated plenum to the high pressure plenum while maintaining theirrespective states. The two energy converters can pump substantiallydifferent flows, with the high pressure energy converter pumping asubstantially lower flow. The system can further include a housing forcontaining the plenums and energy converters, the back of which can havea convex surface curvature compatible with the internal abdominalcavity.

Control systems in accordance with another aspect of the presentinvention can be electrically coupled to one or more electrodes thatsense the heart's electrical activity, and can further comprise anelectronic controller for synchronized release of actuating fluid to thepumping unit for subsequent synchronized evacuation of fluid from thepumping unit. The system can also include a plenum for storage of anon-pressurized volume of fluid of sufficient size to provide a flow atnearly constant pressure during the release interval, said plenum usedfor storage of fluid, a mechanical device or energy converter forperiodically pumping fluid from the storage plenum to the pumping unitand thus attaining a pressurized state in the pumping unit, said energyconverter pumping toward the pumping unit during systole and pumpingfrom the pumping unit during diastole, and a housing for containing theplenums and energy converter, the back of which has a convex surfacecurvature compatible with the internal abdominal cavity.

In an alternative system according to the present invention, the unifiedcontrol system can also include: an internal electronic controller forreceiving both AC and DC supply voltages, an external communicationchannel data stream and generating an actuating signal for the releaseof pressurized fluid, communication channel data stream and internalbattery recharging signals, an actuating means for converting saidactuating signal into a periodic movement of the valve member(s), apumping unit having an associated attachment means wherein the unit isattached directly to the heart or tethered to sites near the heart, saidsites providing a restoring force directed from the apex to the base soas to counter the forces applied to the wrap by the heart, a volumedisplacement chamber containing the energy converter and plenums, ahermetic coupling means for connecting said controller to said energyconverter, said communication channel data streams and internal batteryrecharging signal; a detecting means for generating said actuatingsignal in response to an electrically derived signal from the heartand/or measurement of flows, pressures, tensions related to the heart'sventricles for generating said actuating signal, and a housing the backof which has a surface curvature compatible with the abdominal cavityfor containing said electronic controller, said actuating means, saidblood pumping unit, said hermetic coupling means and said detectingmeans connected so as to form a unified system.

In accordance with another aspect of the present invention, the unifiedsystem can include a rechargeable internal battery for subcutaneousimplantation to supply an internal DC supply voltage. The system canfurther include an external battery for providing a DC voltage to anexternal controller; with the external controller converting DC voltagereceived from the external battery and/or external power supply to ACvoltage for transfer by a subcutaneous energy transformer to power theunified system and/or for recharging external the internal battery. Thesystem can also include a computer interface for connecting said deviceto a computer for control and monitoring of the device; a display meansfor control status and alarm display; a transcutaneous energytransformer for transmitting said AC voltage across the skin; atranscutaneous information telemetry unit for bi-directionaltransmitting said communication channel data streams between saidexternal controller and the implanted components of the system; as wellas a connector for connecting said internal battery to said internalcontroller and an in-line connector for connecting said transcutaneousenergy transformer to the internal battery.

In accordance with another aspect of the present invention, the unifiedsystem can include an implantable stimulator to supply an internal DCstimulus voltage and a rechargeable internal battery controlled by aninternal controller for actuating a selected patient's skeletal musclewith the muscle forming a component of a mechanical pressurizing system.This system can be electrically actuated so as to maintain a desiredsteady-state hydraulic fluid pressure so that pressurized fluid can beheld in reserve in a plenum for transfer to a pumping unit, said pumpingunit powered by the pressurized fluid reserve, and said mechanicalpressurizing system providing for a evacuated side, this side attachedto a second plenum which actively draws fluid from the pumping unitduring diastole.

In accordance with another aspect of the present invention, therechargeable internal battery can be controlled by the internalcontroller for actuating a selected patient's skeletal muscle; with themuscle forming a component of a mechanical pressurizing and electricalenergy generating system. This system can be actuated so as to maintaina desired steady-state hydraulic fluid pressure so that pressurizedfluid can be held in reserve in a plenum for transfer to the pumpingunit via an electrically generating element. The pumping unit can thusbe powered by the pressurized fluid reserve, with this mechanicalpressurizing system providing for a evacuated side, which can beattached to a second plenum that actively draws fluid from the pumpingunit during diastole. The pressurized plenum can further be fitted withan electrically driven valve on the output side and the evacuated plenumfitted with an electrically actuated valve on the input side, and thevalves can be powered by the internal battery and controlled by theinternal controller. The system can further include an optional externalcomputer interface for connecting said device to a computer for controland monitoring of the device; a display means for control status andalarm display; a transcutaneous information telemetry unit forbi-directional transmitting said communication channel data streamsbetween said external controller and the unified system; a connector forconnecting said internal battery to said internal controller and anin-line connector for connecting said telemetry to said internalcontroller.

The control systems of the present invention can be totally implantableand require no physical connections through the skin to the outside. Thesystem can be powered by an implantable battery or directly by atranscutaneous energy transfer system. The battery can be recharged bythe transcutaneous energy transfer system or by a muscle powered device.The device can be controlled and monitored remotely via a transcutaneousinformation telemetry or by said internal electronic controller.

In addition, the device of the present invention can be adapted topneumatic as well as hydraulic activation, and alternative energysources may be utilized such as a Stirling type engine, osmotic pressurevessel, muscle powered generator or a nuclear thermal source.

Anatomical fit to the heart and the maximum contraction length of thewrap are important factors in the clinical success of non-bloodcontacting, volume-amplified inflatable cardiac assist devices. Inaddition to the contraction limitations of the single layer geometry,prior art devices typically also exhibit poor anatomical compliance(‘fit”). In the present invention a superior fit is achieved byproviding a multiple layer geometry with a contraction length largeenough to provide adequate non-contracting regions for attachment andgathering.

Furthermore, the inflation elements are shaped so as to minimize thesize of the non-contracting regions. Attaching a pocket between thepumping unit and natural heart can make refined fit adjustments. Thispocket can be filled with a hydrogel in situ. The gel can form a thinprotective and close fitting layer between the pumping unit and naturalheart. The gel can also be constrained entirely within the pocket orallowed to permeate the pocket so as to form a bond between theepicardium of the natural heart and the pumping unit.

In one preferred embodiment, the major components of the invention areintegrated so that their individual functions are complementary. Thespecifications on these individual components are representative of aunified solution to the problems outlined above and may be unique to thearrangement disclosed. They reflect not only the actuation of aphysiologically acceptable contraction of the heart resulting in enhanceblood flow through the heart, but also relate to issues of maintenance,failure modes, reliability, energy economy, biological compatibility,and quality of life.

Heart assist devices are disclosed to provide an arrangement ofinflation chambers or tubes that, when placed in juxtaposition (e.g.,forming concentric layers of tubes), generate a contraction which issubstantially greater than 36% of the uninflated length orcircumference.

In another preferred embodiment, the inflation chambers can be orientedso that their axes are oriented along, but not necessarily parallel tothe major (longitudinal or vertical) axis of the natural heart. Byemploying elongate tubes that are generally aligned with longitudinalaxis of the heart, efficiency of pumping is further enhanced because thecontractile dimension of the device is aligned with the natural(circumferentially inward) contractile direction of the heart.

The present invention also provides a method of positioning and layeringthe inflation chambers, and an arrangement of non-contracting regions(accounting for as much as 40% of the wrap's circumference) which allowsfor in situ fit adjustments while maintaining a contraction of at least36% over the entire length of the wrap. The present invention furtherprovides specifications for shape and dimension of the inflationchambers, and control of inflation that results in a gentle, effective,and physiologically correct contraction when the wrap is coupled to theheart.

In another aspect of the invention, improved designs and structures forheart-contacting assistance devices are disclosed. For example, thepumping modules can further include a thin, flexible liner between theheart and the inflation chambers to pad the heart. Preferably, the lineris filled with a substance, such as a hydrogel material, which changesthe shape, and not substantially the volume, of the liner whencompressed. This liner provides a customized fit to the natural heart.

In another aspect, the invention also provides a ventricular assistdevice system which can wrap around the natural heart and assist thenatural heart in pumping without coming into contact with blood, andfurther provides a device that mimics the pumping action of the naturalheart and is not directly coupled to the heart, so that in the event offailure, the device does not interfere with the natural pumping actionof the heart.

The present invention also provides a ventricular assist device thatoccupies a volume that is less than the volume of the ejected blood.Likewise, methods are disclosed in which the function of the heartmuscle is assisted by generating an encircling contraction around theheart which exceeds the theoretical limit of single layer balloon wraps,and hence, generates a greater maximum stroke volume.

The pumping devices of the present invention can be fixed with respectto the human heart, but not necessarily directly attached to it, and allof which is free to move with respect to other organs or bones.Moreover, the devices can be configured suitably for use as a right,left, or bi-ventricular assist devices.

The invention can further encompass ventricular assist systems with aportion of the control electronics implanted within the patient and theremainder of the control electronics provided on a small portable unitto be worn on a belt or other clothing or externally attached to thepatient's skin.

In yet another aspect of the invention, multi-layered balloon wraps canbe used in passive assistance devices to provide structures thatrestrain cardiac hypertrophy and mimic the natural resistance of theheart tissue to over-expansion. By choosing an appropriate inflationpressure for the balloon elements and then sealing them, the fluidpressure within the balloons can provide a resistance analogous to theFrank-Starling effect exhibited by cardiac tissue.

In such passive systems, if the heart continues to dilate, the balloonswill flatten more to accommodate the enlargement but the pressureapplied to the heart by the balloons will be greater. Moreover, unlikemesh-type passive girdles, which rely upon an open structure toaccommodate the expansion and contraction of the heart, the multi-layerinflatable structures of present invention permit the use of solid wrapdevices, which are less likely to loss their effectiveness over time dueto tissue in-growth.

In addition, the passive devices of the present invention can beadjusted. For example, if the heart shrinks, the balloon elements can beperiodically filled to a greater extent in order to tighten the wrap andthe pressure applied by the partially inflated balloons will be less.

The invention can also provide architecture that can be adapted to manydifferent geometrical configurations to meet the requirements ofdifferent actuating techniques within the overall constraints of theinvention.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A is a cross-sectional view of a double layer wrap unit cell inaccordance with one embodiment of the present invention;

FIG. 1B is a cross-sectional view of a double layer wrap including theunit cells shown in FIG. 1A showing the geometric relationship of thepartially inflated tubes and the encircled heart represented by adiameter of D_(h);

FIG. 2A is a cross-sectional view of a single layer wrap arranged in aplane shown in a relaxed (diastole) state;

FIG. 2B is a cross-sectional view of a single layer wrap arranged in aplane shown in a partial contraction (mid-systole) state;

FIG. 2C is a cross-sectional view of a single layer wrap arranged in aplane shown in a full contraction (systole) state;

FIG. 3 is a cross-sectional view of a single layer wrap arranged in acylinder showing the systolic and diastolic shapes of the heart musclein a cylindrical geometry;

FIG. 4 is a graphical representation of the relationship between thehydraulic drive pressure to load pressure as a function of tubeinflation;

FIG. 5 is a schematic cross-sectional view of a partially inflateddouble layer device illustrating how unit cells are joined;

FIG. 6 is a further schematic representation of a double layer unit cellillustrating the forces acting on the cell during operation of thedevice;

FIG. 7 is graph illustrating the relationship between pressure ratios,stroke volume amplification and contraction coefficient for a singlelayer wrap;

FIG. 8 is graph illustrating the relationship between pressure ratios,stroke volume amplification and contraction coefficient for a doublelayer wrap;

FIG. 9 is a graph of the contraction coefficient (1−F) as a function offractional volumetric inflation for the single layer and the doublelayer wrap devices, showing that the single layer wrap converges to acontraction coefficient of 0.627 while the double layer wrap convergesto 0.389;

FIG. 10 is graph of performance curves for the single and double layerwraps illustrating regimes of enhanced stroke volume as a function ofpressure;

FIG. 11 is schematic perspective illustration of a pair of tube orballoon elements useful in the present invention;

FIG. 11A is a top view of a single balloon element of FIG. 11;

FIG. 11B is a cross-sectional view of a single balloon element of FIG.11 prior to inflation;

FIG. 11C is a cross-sectional view of a single balloon element of FIG.11 upon inflation;

FIG. 12 is an illustration of the human heart showing the nested sets offiber shells and the spiral path of the fibers in each shell;

FIG. 13 is a schematic illustration of a section of cardiac muscletissue;

FIG. 14 is a further schematic illustration of the varying orientationof cardiac muscle fibers;

FIG. 15 is a partial cross-sectional view of a multi-layer cardiac wrapaccording to the invention;

FIG. 16 is schematic diagram of a united system for cardiac assistanceaccording to the invention;

FIG. 17 is a flow diagram of a mode of operation of the invention; and

FIG. 18 is a schematic diagram of another embodiment of a cardiacassistance system according to the invention.

DETAILED DESCRIPTION

Generally, a ventricular assist system should satisfy a multiplicity ofrequirements beyond the basic requirement of pumping blood. Theintra-thoracic blood-pumping component of the system should not encumberthe natural function of the heart, and preferably has a mean densitycomparable to that of the heart. Moreover, in the event of failure, thedevice should not interfere with the heart. Furthermore, the system lifeand integrity should be sufficient so as to avoid risk of sudden systemfailure. The formation of thrombus within the heart and adjacent vesselsshould be minimized by providing for gentle and physiologically correctpumping action. The device should not damage adjacent tissue ortraumatize adjacent organs by compression or excessive localizedtemperatures. The system should be implantable, preferably withoutconnections through the skin. The system can be supplied periodicallywith an energy source, in some cases that energy may be transmittedtranscutaneously.

First, the geometry of a pumping unit according to the present inventionis described, followed by a discussion of the fit of the inventionaround the human heart. An exemplary construction of the pumping unit ispresented and then the various components of a unified system arediscussed.

Geometry and Construction

FIG. 1A is a top-view of one embodiment of the multi-layer pumpingchamber of the present invention illustrated in a partially-inflatedstate. Pumping chamber 300 includes a plurality of unit cells 302. FIG.1B is a top view of one of the unit cells 302 shown in FIG. 1A. In theillustrative embodiment, each unit cell 302 includes two elongateinflatable chambers or tubes 304A and 304B (collectively and generallyreferred to as tubes 304). One skilled in the relevant art willappreciate that any number of tubes 304 can be utilized in the presentinvention. However, as will be described in detail below, in manyapplications, the two-tube structure 304 is a preferred embodimentbecause of the stability and simplicity that it provided.

The benefits of the multi-layer pumping units of the present inventionmay best be described with reference to a conventional single layerpumping unit such as that described in commonly owned U.S. Pat. No.5,713,954, herein incorporated by reference in its entirety. FIGS. 2A-2Care cross-sectional views of a conventional single layer pumping unit orwrap arranged in a plane and shown in three different states ofcontraction—a relaxed (diastole), partial contraction (mid-systole) andfull contraction (systole) state. Together, these figuresdiagrammatically illustrate the operation of a conventional single layerpumping unit.

The single layer pumping unit 100 includes inflatable elements or tubes102 juxtaposed in a plane 106 to form a flexible wrap. During diastole,the wrap 100 is under tension and not inflated. The cross-section of theindividual tubes 102 is essentially flat as shown in FIG. 2A. The lengthof the wrap 100 attributable to each tube 102 is:$\frac{\pi \quad d}{2},$

and the length of the wrap 100 during diastole, L_(d), may be expressedas shown in Equation (1): $\begin{matrix}{L_{d} = {n\frac{\pi \quad d}{2}}} & (1)\end{matrix}$

where

L_(d)=length of the pumping chamber 100 during diastole;

n=number of tubes 102; and

d=diameter of a fully inflated tube 102.

When the tubes 102 are fully inflated, as illustrated in FIG. 2C, thecross-section of each tube 102 is approximately circular. The length ofthe wrap 100 attributable to each tube 102 is essentially equal to thetube's diameter, d. Accordingly, the length of the wrap 102corresponding to the systolic or contracted configuration, L_(s), may beexpressed as shown in Equation (2):

L _(s) =nd  (2)

where

L_(s)=length of the pumping chamber 100 during systole;

n=number of tubes 102; and

d=diameter of a fully inflated tube 102.

FIG. 3 is a cross-sectional view of a natural human heart and theconventional single layer pumping unit 100 illustrated in FIGS. 2A-2C,showing the cylindrical systolic and diastolic shapes of the pumpingunit 100. As shown in FIG. 3, wrap 100 is joined end-to-end to form acontinuous cylindrical shape. If the tubes 102 are collapsible butnon-distensible, then d is the diameter of the circular cross-section ofan individual tube at any non-negative pressure provided the wrap 100 isnot in tension. When the wrap 100 is under tension, a minimum pressureis required for the tubes 102 to assume a circular cross section with adiameter d. When the tubes 102 are fully inflated, the diameter of thecylindrical wrap 100 corresponding to the systolic state, D_(s.) isapproximated by Equation (3): $\begin{matrix}{D_{s} = \frac{nd}{\pi}} & (3)\end{matrix}$

where

D_(s.)=diameter of the cylindrical wrap 100 corresponding to thesystolic state;

n=number of tubes 102; and

d=diameter of a fully inflated tube 102.

On the other hand, when the tubes 102 are no longer inflated and arecollapsed, then the diameter of the cylinder 200 corresponding to thediastolic state, D_(d,) is approximated by Equation (4): $\begin{matrix}{D_{d} = \frac{n\left( \frac{\pi \quad d}{2} \right)}{\pi}} & (4)\end{matrix}$

where

D_(d,)=diameter of the cylindrical wrap 100 corresponding to thediastolic state;

n=number of tubes 102; and

d=diameter of a fully inflated tube 102.

Since the circumference of a cylinder is linearly proportional to itsdiameter, the diameter of the cylindrical wrap 100 corresponding toboth, the diastolic and systolic states, are reduced by the samepercentage when the tubes 102 are inflated. The percentage of reductionin each is expressed in Equation (5): $\begin{matrix}{\frac{D_{d} - D_{s}}{D_{d}} = {36\%}} & (5)\end{matrix}$

where

D_(d)=diameter of the cylindrical wrap 100 corresponding to thediastolic state; and

D_(s.)=diameter of the cylindrical wrap 100 corresponding to thesystolic state.

This is a limitation on the contractility of both wrap 100 and singlelayer wraps in general. This limitation is imposed by the geometry ofthe wrap 100 and is not affected by the number of tubes 102(n) or thediameter (d) of the tubes 102.

With reference again to FIG. 1B, the tubes 304 are wrapped in anon-distensible sheath 314. The tubes 304 are connected to the sheath314 and to each other at a tangent along their axes, referred to hereinas tangential connection regions 308. The sheath 314 includes two layers316A and 316B on opposing sides of the tubes 304. The two layers 316 areconnected to each other periodically along the length of the sheath 314at cell connection regions 306. The tubes 304 are oriented within theenveloping sheath 314 to form a unit cell 302. A center line 312 residesin a plane that passes through the unit cell 302 and includes the cellconnection regions 306. Assuming the tubes 304 contain substantially thesame volume of fluid or gas, then the plane including the axis 312 alsoincludes the tangential connection region 308 at which the two tubes 304are connected. A second center line 310 shown in FIG. 1B passing throughthe center of the tubes 304 is substantially perpendicular to thecenterline 312 through the center of the associated unit cell 300. Whenthe pumping unit 300 is formed into a cylinder as shown in FIG. 1A, thecenter lines 310 of the tubes 304 are also substantially perpendicularto a tangent 318 on the cylinder 316.

The sheath 314 is preferably configured to fit snugly around the twotubes 304 when the tubes 304 are fully inflated. The mean circumference,C_(s), of the sheath 314 for the exemplary 2-tube layer is approximatedby Equation (6):

C _(s)=2πr+4r  (6)

where C_(s)=circumference of the sheath 314; and r is the fully inflatedtube radius.

The width 320 (FIG. 1A) of the cell 302 is determined by the inflationof the tubes 304. Neglecting tube wall thickness, the collapsed width,W_(c), of a cell 302 as measured along the axis 312 is given in Equation(7). $\begin{matrix}{W_{c} = \frac{C_{s}}{2}} & (7)\end{matrix}$

Since the width of the cell 302 is taken along the axis 312, theinflated width, W_(i), of the cell 302 is given in Equation (8):

W _(i)=2r  (8)

where

W_(c)=the collapsed width of a cell 302 as measured along the axis 312;

W_(i)=the inflated width of a cell 302 as measured along the axis 312;and

r=the radius of a fully inflated tube 302.

For the exemplary 2-layer pumping chamber 300, the fractional lengthchange of the unit cells 302 perpendicular to the tube axis 310 andhorizontal axis 312, F_(d), is given by Equation (9): $\begin{matrix}{F_{d} = {\frac{\frac{C_{s}}{2} - {2r}}{\frac{C_{s}}{2}} = {\frac{\pi}{\pi + 2} = {61.1\%}}}} & (9)\end{matrix}$

where

F_(d)=fractional length change of a 2-tube unit cells 302;

C_(s)=circumference of the sheath 314; and

r=radius of a fully inflated tube 302.

The fractional length change, F_(n), of the circumference, C_(w), of ann-layer wrap is given by Equation (10): $\begin{matrix}{F_{n} = {1 - \frac{2}{\pi + {2\left( {n - 1} \right)}}}} & (10)\end{matrix}$

where

F_(n)=fractional length change of an n-tube unit cell 302;

n=the number of tubes in each unit cell.

Applying these relationships to alternative embodiments of the presentinvention with various number of tubes 304 yields a fractional lengthchange of:

F ₁=36.3%

F ₂=61.1%

F ₃=72.0%

F ₄=78.1%

F ₅=82.0%

where F₁ is the conventional single layer geometry described above.

As shown, the greatest incremental difference in the fractional length,F, is between a single layer wrap geometry (F₁) to a double layer wrapgeometry (F₂). However, significant increases in fraction lengthcontinues to occur as the number of tubes 304 in the unit cells 300increases. When determining whether the double layer embodiment is to beused rather than the alternative embodiments which includes additionaltubes 304, a tradeoff must be made between the utility of incrementalgains in fractional length (F) achieved by the additional layers and theincreased complexity required to prevent the additional tubes frombuckling within the unit cells. (In certain embodiments of the inventionwherein three or more tubes are included in each cell, supportingstructures such as wires are used to secure the tubes 304 in a desiredposition to prevent such buckling).

The associated gains in stroke volume provided by the present inventioncan best be demonstrated on a comparative basis by considering asimplified geometry. If the heart is represented by a cylinder (shown bydashed line 324 in FIG. 1) wrapped by the pumping unit 300, then thestroke volume, S_(v), is given by Equation (11):

$\begin{matrix}{S_{v} = {\left( \frac{C_{wd}^{2} - C_{ws}^{2}}{4\pi} \right)L}} & (11)\end{matrix}$

where

S_(v)=stroke volume;

L=length of the cylinder;

C_(wd)=circumference of the cylinder in diastole; and

C_(ws)=circumference of the cylinder in systole.

For any number (n) of tubes, the change in fractional length, F_(n), ofthe wrap circumference, C_(w), can be defined by equation 12:$\begin{matrix}{{\Delta \quad F_{n}} = \frac{C_{{wd},n} - C_{{ws},n}}{C_{{wd},n}}} & (12)\end{matrix}$

where

ΔF_(n)=change in fractional length (F_(n)) of the wrap circumference(C_(w));

n=number of tubes per unit cell;

C_(wd,n)=circumference of the cylinder with n tubes per unit cell indiastole; and

C_(ws,n)=circumference of the cylinder with n tubes per unit cell insystole.

Rearranging and solving for the circumference of the cylinder in systoleyields the expression given in Equation (13):

C _(ws) ²=(1−F)² C _(wd) ²  (13)

where

F=fractional length of an n-tube unit cell 302;

C_(wd)=circumference of the cylinder with n tubes per unit cell indiastole; and

C_(ws)=circumference of the cylinder with n tubes per unit cell insystole.

The stroke volume, S_(v), follows as shown in Equation (14):$\begin{matrix}{S_{v} = {\left( {1 - \left( {1 - F^{2}} \right)^{2}} \right)\frac{C_{wd}^{2}L}{4\pi}}} & (14)\end{matrix}$

where

S_(v)=stroke volume;

F=fractional length of an n-tube unit cell 302;

L=length of the cylinder; and

C_(wd)=circumference of the cylinder in diastole.

Letting ${L = {1 = \frac{C_{wd}^{2}L}{4\pi}}},$

the stroke volume for cells having n number of tubes is:

S_(v,1)=59.4%

S_(v,2)=84.9%

S_(v,3)=92.1%

S_(v,4)=95.2%

S_(v,5)=96.8%

The diastolic volume (V_(d)) of the heart or interior of pumping chamberor wrap 300 is given by Equation (15): $\begin{matrix}{V_{d} = \frac{C_{wd}^{2}L}{4\pi}} & (15)\end{matrix}$

where

V_(d)=diastolic volume of the heart or interior of pumping chamber 300;

L=length of the cylinder; and

C_(wd)=circumference of the cylinder in diastole.

Conservation of energy provides a fixed relationship between workperformed (W), stroke volume (S_(v)) at a given pressure (P) in the leftventricle, and device inflation volume (V_(H)) and hydraulic drivepressure (P_(H)). Energy conservation is given by the followingwell-known pressure-volume integral equation:

∫P _(H) dV _(H) =∫PdS _(v) =W  (16)

Therefore, there is a tradeoff between changes in inflation volume andinflation pressure (P_(H)), the product of which is a constant.

FIG. 4 is a diagram illustrating a relationship between hydraulic drivepressure (P_(H)) and the load pressure as a function of selected tubeinflation parameter, the tubular inflation angle (θ) introduced above.While the inflation pressure (P_(H)) rises steeply near full inflation(high inflation parameter), the change in inflation volume (V_(tube))changes very little. Conversely, near deflation (low tubular inflationangle), inflation pressure (P_(H)) changes very little, but inflationvolume, represented by tubular inflation angle θ, increases rapidly.These two characteristic behaviors are a consequence of the conservationof energy when a given pressure is applied to an inflatable tube, and isdetermined by the tube diameter (d). By decreasing the tube diameter(d), the inflation volume (V_(tube)) is decreased, which results inincreased inflation pressures (P_(H)). As a result, the desired regimeof inflation pressures (P_(H)) and inflation volumes (V_(tube)) can betailored for any desired layer geometry by varying the tube diameter. Itshould be understood that the maximum attainable fractional contractionper unit length is constant due to the geometry of the unit cell 302.

In the present invention, the double layer design increases strokevolume significantly beyond limits attainable with a conventional singlelayer geometries. The double layer pumping unit 300 is sometimespreferred over embodiments of the present invention having more than twotubes due to the simplicity and inherent stability within the unit cell302 when two tubes 304 are used. The double layer device therefore, inprinciple, offers enhanced design flexibility due to its greater maximumstroke volume. It can provide more pumping work per stroke thanconventional single layer designs and is capable of being adjusted so asto either minimize inflation volume (V_(h)) or minimize the inflationpressure (P_(H)).

Referring to FIGS. 5 and 6, the sheath 314 which forms the unit cell 300in the double layer embodiment does not directly bear on the inflationpressure (P_(H)). In particular, at partial-inflation a portion of thetube circumference is in contact with the sheath causing the walls ofthe tube 304 and the sheath 314 bear the inflation pressure (P_(H)).Therefore, as shown in FIG. 5, forces at the cell connection points 306of two unit cells 300 are directed along the sheath 314. In contrast, ananalogous sheath in conventional single layer designs generates thegreatest contraction when the sheath is in intimate contact with thetube at all stages of inflation. Decoupling the pressure bearingsurfaces between unit cells offers important reliability and longevityadvantages.

FIG. 6 shows a single unit cell 304 of the double layer device 300 in apartially inflated state. An analysis of the performance characteristicsof the double layered wrap 300 is given below. The following assumptionswere used to simplify the analysis:

1) All members are circular and connections between members are tangent.These assumptions are a direct consequence of Laplace's law.

2) There is no friction between the elements, all membrane are stressedand no folds exist.

3) The contact area between the elements is flat.

4) The number of elements is large such that the individual contractionratio is equal to the total contraction ratio.

FIG. 6 is a schematic diagram illustrating the forces applied to onetube 304 of a unit cell 302. As shown in FIG. 6, there is symmetry offorces about a vertical axis. Thus, the following equations may be arederived from well-known force balance relationships for various freebody diagrams within the individual unit cell 302. The vector sum of thetensions in the sheath 314 between adjacent unit cells 304 and the wallstress, T, must balance the force, f, generated by the ventricularpressure in contact with the area between the two tubes where they arejoined. This yields the following relationship shown in Equation (17):

(T+f)Sin(Θ)=P _(H) a  (17)

where

T=tension in the connecting sheath between adjacent unit cells;

f=wall stress in the unit cell sheath;

P_(H)=hydraulic inflation pressure; and

a=contact area.

The vertical component of the tension (T) in the connecting sheathbetween adjacent unit cells 300 must balance the force (f) generated bythe inflation pressure (P_(H)) as follows:

T Sin(Θ)=P_(H) w  (18)

where

w=width of wrap or length of balloon;

T=tension in the connecting sheath between adjacent unit cells; and

P_(H)=hydraulic inflation pressure.

The theoretical performance of a conventional single wrap design anddouble layered pumping device according the present invention have beenanalyzed. FIGS. 7 and 8 are graphical representations of the hydraulicamplification and operating pressures as functions of devicecontraction. Device contraction is represented by a contractioncoefficient. A contraction coefficient of 0 indicates no contraction.FIG. 7 (representing a conventional single layer arrangement) shows thata single wrap operating at an end systole pressure or maximum inflationpressure of 10 times the ventricular pressure(P_(H)) (a pressure ratioof 10), will develop a hydraulic amplification of 3.5 when thecontraction coefficient is 0.2 .

As shown in FIG. 8, a double layer device operating under similarconditions will generate a hydraulic amplification of 2.5. Hence, thefluid requirements of a conventional single layer device are lower thanthose of the double layer device of the present invention. Since theenergy requirement is equal in both cases, the single layer device mustoperate at a higher average pressure to achieve the same end pressure.

FIG. 9 provides another illustration of the relationship between thecontraction coefficient and a given inflation parameter for single anddouble layer wraps. Volumetric amplification A_(v) is the ratio of thestroke volume of the heart S_(v) divided by the change in balloon volumeduring a stroke. That is, $\begin{matrix}{A_{v} = \frac{S_{v}}{\Delta \quad V_{balloon}}} & (19)\end{matrix}$

FIG. 10 is an illustration of the relationship between the pressureapplied to an individual tube element 304 and stroke volume for singleand double layer wraps.

These graphic results illustrate two important features thatdifferentiate the present invention from conventional single layerdevices. In the case of a conventional single layer pumping unit, theratio between stroke volume (S_(v)) and change in balloon volume isgreater at every point during the contraction cycle than in the doublelayer device, regardless of the chosen contraction coefficient orinflation pressure. This is a direct consequence of the intrinsicdifferences between the PV curves of the two devices.

In the case of the double layer wrap 300, the maximum contractioncoefficient is always larger than for the single layer wrap (see FIGS. 7and 8). This characteristic is independent of the unit cell size. If anend contraction distance is specified, then the double wrap reaches thisendpoint at a lower end inflation pressure for contraction coefficientsless than 0.75.

In summary, the multiple-layer pumping chambers of the present inventionhave unique geometrical configurations that enable them to develop largecontractions at low end inflation pressures.

In addition to the nondistensible sheath 304 in which unit cells 302 areformed to hold the elongate tubes 304, the pumping unit 300 can alsoinclude an inner pocket for in situ fit adjustment.

The sheath 314 can be composed of two layers of a woven, monofilament,polyester sheet material. For example, the filament size can be 30-100microns, and the open area can be about 25% to about 50%. A pattern canbe transferred to the fabric, and the pockets formed using either atight sailing-type stitch or a polyester bonding technique.Alternatively, the sheath can be formed from a single urethane solutioncast. The mandrel for casting a polymeric sheath can be a low meltingpoint material or soluble polymer such as PEG, with through-holes thatprovide interconnections. The mandrel is dipped in a flocked urethanesolution. The mandrel is removed by melting or dissolving. The resultingsheath is nondistensible yet flexible, and the pockets are delineated bythe interconnects formed at the through-holes.

FIG. 11 is schematic perspective illustration of a pair of tube orballoon elements useful in the present invention. FIG. 11A is a top viewof a single balloon element. Each of the balloon tubes 304A, 304B has aflatten cross-section, as shown in FIG. 11B prior to inflation and acylindrical cross-section upon inflation as shown in FIG. 11C. Theballoon tubes preferably have rounded ends 301 and 303. The balloonpairs can be made as a single unit or as two distinct balloons, asshown, and coupled to an inflation source via coupler 305. The balloonsare preferably made of polyetherurethane. Mandrels of the desired shapeare dipped to form balloons either as a pair or as singles. Thedesirable wall thickness being uniform and in the range of about 2 toabout 10 mil.

Heart Fit

The manner in which the pumping unit 300 is secured or fit to a humanheart is now described. FIG. 12 is a partial cut-away view of the humanheart showing nested sets of fiber shells with each such shell having aspiral path of fibers. The shape of the heart is defined by theepicardium following a geodesic contour of a surface such that there isminimum surface; in particular, this shape is commonly referred to as aparaboloid of revolution. As shown in FIGS. 12-14, the myocardium can bevisualized as a nested set of fiber shells. These figures have beenderived from those provided in D D Streeter, C Ramon, “Muscle PathwayGeometry in the Heart Wall,” IEEE Journal of Biomechanical Engineering,105:367-371 (1983); D D Streeter, W T Hanna, “Engineering Mechanics forSuccessive States in Canine Left Ventricular Myocardium,” Circ.Research, 33:656-664 (1973); D D Streeter, R N Vaishnav, D J Patel, H MSpotnitz, J Ross, E H Sonnenblick, “Stress Distribution in the CanineLeft Ventricle During Diastole and Systole,” Biomedical Journal,10:345-363 (1970); and M A Ross, D D Streeter, “NonuniformSubendocardial Fiber Orientation in the Normal Macaque Left Ventricle,”European Journal of Cardiology, 3(3):229-247 (1975), all of whichdescribe this construction of the human heart and which are herebyincorporated by reference in their entirety.

On each of the shells 1202 the muscle fiber 1204 travels in a spiral.Adjacent fibers 1204 in the same shell 1202 are substantially parallelas shown in FIG. 12. Proceeding from shell-to-shell, the angle of thefibers 1204 changes gradually. That is, the fibers 1204 in one shell1202 are not parallel with the fibers 1204 in an adjacent shell 1202;instead, they are offset from each other by a few degrees.

Referring to FIGS. 13 and 14, the relationship between the shells 1202of the heart are now described. A block 1302, 1304 removed from the fullwidth of the wall 1300 of the heart 1202 reveals that the fiberorientation resembles that of the open blades of a Japanese fan, asshown by reference numeral 1306. On each shell 1202A, 1202B, etc.,represented by a blade of the fan, the fibers follow the line of theblade, but each fiber is randomly tethered to its neighbor within theblade and from blade to blade. The fibers are restricted to a lateralsliding or rolling motion, and no fiber acts independently. Generallyspeaking, left ventricular fiber angles (in diastole and systole) can beseen as a function of percent wall thickness. The mean fiber angle(spatial orientation) changes continuously from about +60 degrees on the-endocardium surface to about −60 degrees on the epicardium surface.

Generally speaking, a two dimensional sheet structure can be bent into aparaboloid of revolution if V-shaped sections are periodically excisedfrom the sheet, and adjacent edges thus formed are joined. However, ifthe V-shaped sections lie on a line, the resulting structure will be acone. A cone can be distinguished from a parabola in that the edge ofthe parabola has an increasing slope. If the V-shaped sections are cuton two arcs such that when joined together along with the wrap ends, ashape which has a small slope on the apex side and a large slope of thebase side is formed. The curvatures of the arcs are such that the arcforming the apex side has a shorter length than the one forming the baseside is formed. Although the V-shaped sections forming the apex and thebase are paired in one-to-one correspondence, the distance betweenadjacent V-shapes on the apical side is shorter than that on the baseside. Based on these geometric principles, a pumping unit can beconstructed in this way such that it develops a circumferentialcontraction around the heart when the tubes are oriented with their axisrunning from apex to base. Accordingly, the tubes can be located betweenapical-base paired V-sections. An enclosing sheath can be constructedfrom two such sheets, where pockets are formed by bonding or stitchingthe sheets together to form pockets. The outlines of the pockets wouldfollow the contours of the V-shaped sections. Preferably, these pocketsare slanted with respect to the axis of the heart so that they effect acontraction that is approximately 30-60 degrees from the midline.

Thus, the tubes (e.g., inflatable balloons) can be mounted in thepockets, such that the open end of the balloon extends out the apicalor, alternatively the base side of the pumping unit. The portion of theballoon that is exposed is preferably fitted with a device forconnection to a distribution network. The balloons are preferablytapered at both ends such that when they are inflated, a fabric sheathcan be brought tightly against the balloon pair at all points. This“wrap,” when placed around the heart, affects ejection of blood byreducing the circumference of the heart by direct volume displacementdue to distention of the balloons toward the heart, and by a pressingaction. In one preferred embodiment, the balloons can transform from aparabolic curvature induced by the diastolic heart to a straight tubeconfiguration as a consequence of inflation and thereby eject greatervolumes of blood.

In some applications it may be preferred that both ends of the balloonbe closed. In such an embodiment, the connection point to the fluidsource can be located between the ends and directed away from the heartso that any connecting tubing does not interfere with circumferentialcontraction. However, in tubes that are tapered at both ends, theweakest point of attachment is midway between the ends. The stresses canbe defined as follows:

This choice is a consequence of considering the stresses at point r,σ(r), in the free balloon wall: $\begin{matrix}{{\sigma (r)} = \frac{pr}{t}} & (20)\end{matrix}$

where:

p is the pressure inside the balloon;

r is the radius of the balloon; and

t is the thickness of the balloon wall.

Consequently the highest local stresses are at points of largest radiusof curvature. For this reason, a preferred embodiment entails taperedballoons with one end closed with a hemisphere, the hemisphere radiusbeing half the balloon diameter at the point where the hemisphere joinsthe taper.

Considering the unit cell 300 described above comprising the twoballoons 304A and 304B and the sheath 302. The inflated length of thesheath—l_(i)—may be represented by the expression given by:$\begin{matrix}{l_{i} = {l_{o} + \frac{d}{2}}} & (21)\end{matrix}$

where:

l_(i) is the inflated length of the sheath;

l_(o) is the part of the sheath corresponding to the straight part ofthe balloons 304; and

d is the inflated diameter of the balloon.

The flattened or deflated length of the sheath, l_(d,) is given by:$\begin{matrix}{l_{d} = {l_{o} + \frac{\left( {\frac{\pi}{2} + 1} \right)d}{2}}} & (22)\end{matrix}$

The shrink ratio, R, defined as the change in linear dimensions relativeto the original linear $\begin{matrix}{R = {\frac{\left\lbrack \frac{l_{o} + {\left( {\pi + 1} \right)d}}{2} \right\rbrack - \left\lbrack {l_{o} + \frac{d}{2}} \right\rbrack}{\frac{l_{o} + {\left( {\frac{\pi}{2} + 1} \right)d}}{2}} = \frac{\pi}{\frac{4l_{o}}{d} + \pi + 2}}} & (23)\end{matrix}$

dimension l_(d) is:

The shrinkage, R, is a function of the ratio between the diameter d andthe length l_(o). This ratio is valid for any number of unit cellsjuxtaposed provided l_(o)/d is identical for all such unit cells.

In one preferred embodiment, to achieve a desired fit, the wrap 300 isconfigured so as to shrink in proportion to the natural apicalcontraction of the heart. Typically, this is approximately 12%. LettingR=0.12 yields a length to diameter ratio of l_(o)/d=5.26. The diameterused here is the diameter where the balloon merges to form a hemisphere.

This does not take into consideration the taper in the balloon 304.There are a variety of tapers which are possible. All tapers involve acontinuous transformation from a maximum diameter to a minimum diameter.The exact mathematical formula used to determine this profile willaffect R as expressed below: $\begin{matrix}{R = \frac{\frac{\left\lbrack {l_{o} + {\left( {\frac{\pi}{2} + 1} \right)d\quad \lambda}} \right\rbrack}{2} - \left\lbrack {l_{o} + \frac{d}{2}} \right\rbrack}{\frac{l_{o} + {\left( {\frac{\pi}{2} + 1} \right)d\quad \lambda}}{2}}} & (24)\end{matrix}$

where the large diameter is D=dλ. The taper then is expressed by λ. Inthe particular case where D=10 mm, R=0.12, and l_(o)=80 mm, the valuefor lambda is λ=2.92. This value for lambda is compatible with therequirement that the taper provide a good fit to the heart when theballoon is deflated. It should be noted that to attain a 12%longitudinal contraction in the single layer design with tubes of thesame diameter, two wraps would be required to cover the heart (asindicated in Ser. No. 08/490,080). In this event, a tapered designemploying two wraps would impede ventricular filling at the junction ofthe two wraps.

Although a parabolic profile is a good fit to a side projection of theheart onto a plane, there are numerous variations in the longitudinaldirection which translate into various distances from the heart's axiswhen measured on a line traveling in the circumferential direction onthe epicardium. When the wrap performs a pumping action on a deformable“corrugated” surface such as that of the heart, the corrugated regionsare filled by extension of the balloons and contraction of the wrap wheninflated. This represents lost stroke volume, and studies on casts ofdog hearts reveals this loss to be approximately 100% of the realizedstroke volume.

The preferred embodiment contains a segmented or single pocket whichextends over the entire surface of the wrap on the side facing theheart. This layer is preferably be filled with a compliant materialwhich satisfies the following criteria:

1. have approximately the specific gravity of water

2. have the viscosity of water when first introduced into the wrap sothat an in situ fit can be made

3. have the properties of a solid after a fit is made, such that thematerial does not pool to the apex and inhibit the proper filling of theheart

4. remain unchanged in a wet, 37° C. environment

5. be easily deformable but not change appreciably its volume

6. be nontoxic and biocompatible

7. does not release a gas, or is excessively exothermic or endothermicwhen transforming from a liquid state to a solid state

8. possess a low durometer reading so that it does not excessively loadthe wrap

9. possess an elastic quality which will aid in ventricular filling, andnot impede ventricular filling in the event of system failure

10. have a slow setting time relative to the beat frequency of the heartso that an average heart shape results

These requirements are satisfied by a number of generally availablehydrogels. One example of a hydrogel which may be used in conjunctionwith the present invention is Hypol™ PreMA G series manufactured byHampshire Chemical Corp., Lexington, Mass. Hypol forms a polyurethanematrix hydrogen bonded to up to 95% water by weight. The formation of ahydrogel layer between the wrap and the heart is an important aspect ofthe present invention. One embodiment provides for two entrance portsand an exit port affixed to the pocket. Preferably, the port isdetachable and the entrance ports are positioned at opposite ends of thewrap with the exit port at a location midway between the two exit ports.

In one embodiment, a mixture of 10% Hypol and 90% water is mixed in asealed blender type apparatus, or alternatively a mixing syringe. Theingredients mixable by breaking a separating membrane. The resultantmixed solution introduced to the wrap through entrance ports. The pocketis filled until fluid is detected at the exit port, and subsequently theexit port is scaled and the pocket gently pressurized so as tocompletely fill all voids between the wrap and the beating heart. Thesolution sets in 60 seconds, allowing for the liquid state of the pocketto be in contact with the beating heart for approximately 30 seconds.The solution sets in a continuous fashion, reaching a maximum durometerreading in about 10 seconds after setting begins. The setting timeallows for the pocket to be modified by changes in the hearts shapethrough at least 10 heart cycles, so that the final shape is a fit tothe mean heart shape. This provides for the least loading on the wrap.

Although both the single and double layer designs generate a radiallypressing action, the following analysis shows that the pressing actionis negligible for the single layer design. This action should be viewedas distinct from a direct volume displacement, which also occurs, and iscaused when inflation of the balloons creates gaps between them.Referring now to, FIG. 15, the dashed line, or centerline, defines thelocus of points which contract under inflation of the balloons. Aninscribed line defines the surface of the heart. In the single layerdesign the gap is small enough for the heart or hydrogel to conform tothe gently corrugated surface. In the double layer design the gap isdeeper. Assuming now the heart does not conform to the device surface,but instead lies on the inscribed line, it is possible to calculate ahydraulic gain factor specific to the gap forming phenomenon. In thiscalculation only the inflation volume between the centerline and thenatural heart is considered. For the single layer design the inflationvolume per unit cell is one half the volume of a single balloon, and forthe double layer it is one balloon volume. For the sake of comparisonone can assume a spacing between adjacent balloons in the single layerdesign of 1 inch. The spacing between adjacent balloons in the doublelayer design can be 1.5 inches. Assume an inflation pressure of 30 psi.At this pressure the single layer design develops a maximum contractionof 34%, whereas in the double layer design it is only 40%. Recall thatthe theoretical maximum at any pressure for the single layer is 36%,whereas in the double layer design it is 61%.

The gap formed has a cross section defined by:

A=S×C×h  (25)

where:

S is the spacing;

C is the contraction; and

h the displaced distance from centerline.

The displaced distance from the centerline, h, can be approximated byassuming

(cx)² +h ² =H ²  (26)

where:

x is equivalent to S/2, and the hypotenuse H is S/2.

Then

h _(sigle)={square root over ((1−C ²)x)}=0.37S  (27)

and h_(sigle)=0.4oS.

The gap areas for a large balloon is 0.24 (2.2) for a single layer, and0.54 (2.5) for a double layer. The portion of the gap filled by thecross section of the balloon is 0.11 for the single layer device and0.22 for the double layer device. The gain factors are given inparentheses above. It is important to note that if the double layer werefully inflated to 59% contraction of a possible 61%, the gain factor isconsiderable larger for the double layer design: 0.84 (3.8).

Attachment

The pumping unit of the present invention may be attached to the heart.There are a variety of attachment approaches that are discussed in theliterature. These approaches can be grouped into 6 categories: 1) directsuture of the wrap to the heart such as that described in U.S. Pat. No.5,098,369; 2) those that use a strap through the transverse sinus ormajor vessels which are attached to the wrap, such as those described inU.S. Pat. Nos. 5,131,905 and 4,536,893; 3) those that utilize variouschest cavity attachments, including attachment to the sternum or the ribcage, such as those described in European Patent No. 0 583 012 A1 andU.S. Pat. No. 4,925,443; (4) vacuum between the wrap and the heart, suchas described in U.S. Pat. No. 5,119,804; (5) those that use a staple toattach the wrap to a surrounding tissue such as the pericardium,described on U.S. Pat. No. 4,690,134, or the diaphragm and (6) the useof mechanical griping elements on the inside of the wrap as described incommonly owned, co-pending U.S. patent application Ser. No. 09/661,885by Milbocker filed Sep. 14, 2000, the teachings of all of thesereferences are incorporated herein. Other attachment mechanisms willalso be apparent to those skilled in the art.

In one preferred embodiment, the heart is secured so as to beindependent of the skeletal system in the event trauma should occur tothe skeletal attachment. Such a trauma would result in impaired heartpumping, even if said trauma would not have otherwise affected theheart. Furthermore, as Heilman et al teach in U.S. Pat. No. 4,925,443,it is necessary to provide for a mechanism to follow the naturalmovement of the heart, especially torquing forces produced by thecontracting heart. In the case of direct attachment to the heart, eitherthrough suture or bands placed on the vessels or sinus of the heart,produces localized points of abrasion and force centers which may resultin long-term thrombus production or failure of the tissue. One possibleacceptable direct attachment to the heart would involve promotion ofingrowth between the heart and wrap. This is a typical conditionobserved in successful cardiomyoplasty, where the latissimus dorsi bondsto the epicardium. The vacuum attachment method is not preferred sinceit requires an external vacuum source. Attachment to surrounding tissueis the preferred attachment method of the present invention. Thesetissues can be used in combination, such as posterior attachments to thepericardium and apical attachment to the diaphragm.

When the wrap is placed around the natural heart and the filling andevacuation of the pumping unit is in phase with the systole and diastoleof the natural heart, then the contraction of the wrap around theventricles of the heart decreases the circumscribed diameter therebycausing the ventricles to eject blood. The wrap operates equally well asa single ventricular assist, however, the wrap ends must be sutured tothe septal line of the epicardium. The contractility, and thus theejection fraction, of the wrap is independent of the number of tubes orthe heart dimension. The ejection fraction is only a function of thehydraulic pressure.

FIG. 16 illustrates a unified system according to the inventionimplanted in the human body. In the illustrated system the pumping unit1 is shown as a cuff placed around the ventricles of the natural heart2. Hydraulic fluid is used to drive the pumping unit, as also taught inU.S. Pat. No. 5,713,954 issued Sep. 3, 1998 entitled “Extra-CardiacVentricular Assist Device,” herein incorporated by reference. Hydraulicfluid is preferred for the following reasons: 1) a pneumatic system canbe susceptible to fluid permeation through the balloons and connectingtubing, whereas the proper choice of a driving fluid substantiallyeliminates this problem, 2) a gas pressurized system results in a volumeexpansion in the event of a ruptured line or balloon, whereas a fluid isincompressible, 3) in the event of system failure, the fluid can bedrained out of the balloon, and 4) the volume of the pressure and vacuumplenums is smaller. The hydraulic fluid is pressurized by energyconverter 3 and fed directly to a spring-loaded hydraulic reservoir 4.The hydraulic reservoir 4 supplies pressurized fluid to the pumping unit1. Fluid flows from the pumping unit 1 to a second spring-loadedhydraulic reservoir 5. This reservoir 5 is connected directly to theenergy converter 3 to form a closed loop. The energy converter 3, theoutput of reservoir 4 and the input of reservoir 5 are electricallycontrolled by virtue of internal electric controller 6 which is poweredby an internal battery 7. When the internal electric-controller 6 alsosupplies the driving energy for the energy converter 3, it is coupledalso to external battery 7 via a transcutaneous electrical terminal(TET) 8. For additional details on transcutaneous energy transfer andtelemetry, see commonly-owned, co-pending U.S. patent application Ser.No. 09/304,198 filed May 3, 1999 by Kung, the teachings of which areincorporated by reference.

The energy converter 3 may consist of a hydraulic pump coupled to abrushless electric motor to shuttle fluid between reservoir 5 andreservoir 4. This arrangement allows for unidirectional, and continuousoperation of the electric motor. Activation of the outlet valve 9 onreservoir 4 and the inlet valve 10 on reservoir 5 is synchronized by acontrol signal generated by detection of the R wave from the ECG signalcollected by electrode 11. Continuous adjustment of the hydraulic pumpoutput allows the level of cardiac assist to be varied on a movingaverage basis. The duration of the valve opening on reservoir 4 allowsthe level of assist to be varied on a beat-by-beat basis. Since it isthe tendency of the heart to compensate a long beat interval with ashort beat interval, this arrangement insulates the motor from erraticvariation in beat rate. The motor pump rate will respond to meanpressure variations in the reservoir 4.

Alternatively the energy converter may be a muscle driven unit as taughtin U.S. Pat. No. 5,344,385. In this case the muscle used is thelatissimus dorsi, and may be stimulated by a low current signal from theelectronic controller 6. This approach has the distinct advantage ofdirectly converting mechanical energy to hydraulic energy. The otherapproach has losses at: 1) the external battery, 2) the TET, 3) thewindings in the motor, 4) the conversion of mechanical energy in themotor to hydraulic energy. Typical power consumption for such systems isapproximately 20 W. The human heart produces about 1-2 W of hydraulicpower. With nearly 100% efficiency between the muscle attachment and thehydraulic compressor mechanism, the power required of the muscle isapproximately 2 W. The power available from skeletal muscle has beenestimated to be 3 to 15 mW per gram (Geddes et al., (1991), Trans.ASAIO, 37:19-23), a muscle mass of about 116 to 350 grams is generallysufficient within the scope of the present invention. The latissimusdorsi, pectoralis, psoas major, and rectus abdominous muscles are majorskeletal muscles capable of supplying sufficient power output. A deviceas taught in U.S. Pat. No. 5,344,385 makes best use of the muscle byemploying the muscle in situ, with a normal pre-stretch, and pulls intension rather than squeezes, as is the case when the muscle is wrappedaround the heart. The time required for contraction is also important.The muscle in the present invention would not be directly coupled to theheart. For beat rates of 70 to 80 beats per minute, the systolic anddiastolic intervals are 300 to 500 ms. A stroke length of 4 cm with acontraction time of 300 ms would require an average velocity of 13.3cm/s. However, 13.3 cm/s is a fast contraction for any skeletal muscle,and is particularly fast for a large muscle. A shorter stroke lengthresults in lower contraction velocities. Thus the optimum design wouldfavor lower contraction velocities and higher exerted forces.

The natural heart pumps blood primarily through circumferentialcontraction. Most of the diastolic to systolic volume change is derivedprimarily from the 20% change in the circumference and to a lesserextent to a 9% change in the axial length. The volumetric change in thevolume enclosed by the epicardial surface is 36% from the relaxed(diastole) position to the fully contracted (systole) position. Asdescribed earlier, for a two-layer geometry the pumping unit generatesan 85% ejection fraction. Although a cylindrical geometry was used inthe above calculation, when the unit cells are appropriately tapered,the contraction as a percentage of the local heart circumference is thesame at all points on the wrap.

With this hydraulic design, the natural heart having a typicalmyocardium thickness, a heart base diameter of 80 mm and an axialventricular length from apex to base of 50 mm, a left ventricular wrapresults in a stroke volume of 118 ml. This stroke volume exceeds thatexpected from a normally operating left ventricle. Since the drivepressure rises exponentially near full inflation, normal stroke volumesare achieved at lower inflation pressures.

The pumping unit 1 is activated by a signal from an implanted epicardiallead in a myocardial region of the natural heart not in contact with thepumping unit. Suitable locations are in the apical region or near theright atrial appendage. Release of hydraulic fluid to the wrap is to betimed with the R wave produced on this lead. Detection of the R wave canbe accomplished with techniques in use in implantable defibrillators.The systolic duration is preprogrammed (as taught in U.S. Pat. No.5,713,954 issued Sep. 3, 1998 entitled “Extra-Cardiac Ventricular AssistDevice,” herein incorporated by reference) to match its functionaldependence on the beat rate (BR) in beats per minute expressed as

t _(S)=549−2BR.  (28)

The time delay t_(R) is the delay between pump initiation relative tothe R wave, that is, the interval when the pumping unit is in anevacuated state. The exact coincidence of the start of the diastolicduration and the ECG T-wave is not critical. Irregular rhythms in thenatural heart will trigger the system to assume a no pump status. Inthis state, the pumping unit is evacuated. In the event the ECG signalis lost or too noisy to detect the R-wave, the device could begincontracting at beat rates consistent with maintaining physiologicalfilling pressures. Thus the Unified system would provide the userprotection against sudden physiological electrical as well as mechanicalfailures which may occur in the diseased natural heart.

The control scheme is motivated by the intention to operate the systemsuch that the natural heart maintains a full systolic stroke in everybeat. A given hydraulic stroke volume corresponds to an uniquecontraction; and therefore, there is a one-to-one correspondence betweenhydraulic stroke volume and physiological stroke volume. If at a givendrive pressure the hydraulic stroke volume does not reach the programmedtarget value, the power required to achieve full stroke can will beautomatically adjusted, on a beat by beat basis if necessary. Duringdiastole, the hydraulic pressure will be measured to provide anindication of the end diastolic pressures. From this information abaseline beat rate may be determined, which will automatically be thedefault operating parameter for the system if the heart should stopbeating.

Synchronization is achieved by sensing the natural rhythm of the heartor through implantable pacer electrodes. As is taught in U.S. Pat. No.5,713,954 issued Sep. 3, 1998 entitled “Extra-Cardiac Ventricular AssistDevice,” herein incorporated by reference, and is well known in the art,two basic approaches are available. The choice between using the P-waveor the R-wave of the ECG signal is determined by the conduction fitnessof the heart.

It is well known that mechanical disruption of the heart can also createelectrical changes in the heart. The P-wave may be preferred in thisrespect since the right atrium is free from mechanical contact with thepumping unit. This approach assumes the right atrium is electricallysound, and free of frequent atrial flutter or fibrillation. Clearly AVblock would exclude use of the P-wave. Atrial sensing is accomplishedwith a lead sutured to the right atrial appendage. Alternatively, in thecase of R-wave sensing, a corkscrew electrode is attached to theventricle near the apex at a location free of wrap contact. In theR-wave sensing mode, it is possible to combine sensing with a means forattaching the pumping unit to the heart. Bipolar electrode designs areknown which provide for localized signal reception of the P-wave,providing a reduced noise signal. Unipolar leads are sufficient forR-wave sensing since this is the simplest type of electrode forventricular epicardial fixation.

Modifications to the control algorithm to accommodate either signaldetection method are straight forward and ensure synchronous contractionin the pumping unit. Typically P-wave sensing requires the systolicinterval to begin 160 ms after P-wave detection. The systolic intervalcan be initiated immediately upon detection of the R-wave. U.S. Pat. No.5,713,954 teaches a possible algorithm based on the prior R-R interval.

U.S. Pat. No. 5,713,954 teaches of a subdermal port for draining thepumping unit in the event of failure of the hydraulic pumping capacity.Such a subdermal port can be accessed through a skin puncture with anarray of 15 gauge needles. The procedure would involve extraction ofhydraulic fluid using a syringe. The result would be collapse of thewrap. An automated or manually driven pneumatic pump could then beconnected to the needle manifold to substitute for the failed hydraulicpumping unit. Pneumatic extra-corporeal activation is preferred, sincethe pressure drop for a hydraulic fluid through the needle array wouldbe unacceptably high.

Implantation of the unified system can require a median stomatomy. Anappropriately sized pumping unit would be placed on the natural heart.Laces or adjustable clasps on the V-shaped sections would be used toachieve a closer fit. The hydrogel pocket would be filled as describeearlier to provide a custom fit. The pumping unit would be anchored tothe epicardium either by suture or by an adhesive applied between theepicardium and the wrap. The energy converter and the hydraulicreservoirs would be implanted as a unit in the thorax. The energyconverter unit would be anchored to the rib cage with a flexiblehydraulic connection to the pumping unit and an electrical cabletunneled through the costal diaphragmatic region to the electroniccontroller that could be implanted in the abdomen. One preferredembodiment of the present invention combines the capacity to rununassisted on the patient's own muscular power as well as allowingassistance through a transcutaneous energy transformer.

With reference again to FIG. 16, a unified system can include theimplanted components: a pumping unit, and electronic controller,internal battery, and transcutaneous energy collection and telemetrysystem. It further includes external components: a transcutaneous energytransmitter and telemetry system, computerized data collection,analysis, and encryption system; and associated batteries, electricaladapters to conventional electrical energy sources, and gear for holdingthese components.

Actuators 9 and 10 converts the actuating signal from the electroniccontroller 6 into a periodic pressure/volume wavefront in the pumpingunit 1. This generates a rhythmic displacement of fluid in the pumpingunit, causing a rhythmic fill and drain of fluid in the plenums 5 and 4.The actuating signal is derived from a physiologic ECG signal obtainedfrom an electrode 11 placed on the heart. The energy converter 3 can bean electrical motor powered by the implanted battery 7 or muscle poweredor a combination of both. In the case of the muscle powered energyconverter, the muscle receives a stimulating signal from the internalelectronic controller. This signal can be in response to time varyingpumping requirements or to physiologic requirements of the actuatingmuscle. FIG. 17 provides the decision tree that can be used to governeither mode of operation.

The overall system provides for the optional use of the TET system topower the device. The decision can be sent via the external telemetry tothe internal controller, or can be made, as illustrated, predicate onvarious conditions within the implanted system. In either mode ofoperation, mean pressures are maintained in the plenums in response tomean pumping demands and desired contraction requirements.

The components of a hybrid energy converter are illustrated in FIG. 18including internal controller 116, internal rechargeable battery 105,internal telemetry transceiver 119, energy converter 110 and plenum 117,as well as external computer 121, external telemetry transceiver 118 andexternal energy charging source 103. The device is designed to provide aquasi-continuous output.

The internal battery 105 can be connected to the internal controller116, telemetry transceiver 119, and energy converter 110, via hermeticsealed electrical connectors. The internal battery 105 is preferablyimplanted subcutaneously in the abdomen. In the preferred embodiment theinternal battery package uses rectangular prismatic nickel/cadmiumcells, however other battery chemistries could be utilized. These cellsare housed in a custom designed laser welded titanium enclosure with ahermetic connector suitable for implantation. The battery housing isarched to provide for good anatomical fit. Preferably the hermeticconnector is an internal connector, used for allowing easy replacementof the battery package. In this way, the conductor set need not bechanged when and if the internal battery is replaced. The electricalconductor set contains electrical connections for the supply of DCvoltage to the internal electronic controller, for the recharging of theinternal battery and for activation of -an-audible warning alarm housedin the battery enclosure.

The internal TET/Telemetry unit 105 and 119 can also be implantedsubclavicularly. The internal electronic controller 116 receives thetransmitted external TET/Telemetry signals from internal TET/Telemetryunit 119. Power signals are passed to a power regulator and telemetrysignals to a telemetry subsystem of the internal controller viaconductors. A power regulator can generate the DC voltage for supplyingthe circuits of the internal controller 116 and charging battery 105.The telemetry subsystem exchanges signals with the internalTET/Telemetry unit 9 and with control unit 14. A sensor (as describedabove) can monitor the pressure in pressure plenum 117 and/or hydraulicfluid consumption. In one embodiment the sensor can be an electrodeattached to the heart and monitors the electrical activity of the heart.This allows the electronic controller 116 to provide a Starling typeresponse to physiological changes in oxygen demand. The internalcontroller 116 processes the data received from the detection means 19and 16 and instructs the motor commutator 113 or the muscle 120 togenerate pressure levels and instructs the inlet/outlet valves to adjustbeat frequency, systolic and diastolic duration. Internal controller 116exchanges information with external computer 121 through telemetrycircuit 118.

The transcutaneous energy transformer (TET) and transcutaneousinformation telemetry (Telemetry) systems can uses wire coils toelectromagnetically couple power into the body without perforation ofthe skin. The Telemetry uses infrared components embedded in theinternal and external TET/Telemetry units to transfer the communicationchannel data streams into and out of the body without EM interference.The external TET/Telemetry system is connected to the external computervia a conductor.

The external computer 121 can contains a portion of the circuitry forTET/Telemetry systems as well as algorithms which can diagnose andspecify solutions to emergency situations. In particular the externalcomputer contains detailed specifications for when to deactivate thesystem so that the pumping unit does not interfere with normal heartfunction. Additionally, the external computer contains specificationsfor determining when to operate in a total heart assist mode, usuallyindicated when abnormal or no electrical signals are received from theheart. It also allows the internal controller to operate in asynchronous and asynchronous mode. The external computer also producesand receives the communication channel data signal for control andmonitoring of the implanted system and generates recharging signals forrecharging the external battery at predetermined intervals. The externalelectronics can be a compact unit adapted to be worn on a belt.

Various factors contribute to the fit of an extra-cardiac assist device,among these are: balloon shape (taper), pocket shape, overall shape ofthe sheath, balloon wall thickness, sheath material stiffness, means forjoining the ends of the device, proportion of non-contracting space tocontracting space, orientation of the balloons, and size of theballoons. Some design considerations oppose one another. For instance,in the single layer design the circumference of the sheath must be asclose as possible to the circumference of the balloon, yet balloon wallthickness and stiffness require the circumference of the sheath to belarger if the balloon is to collapse without wrinkles. Furthermore,overall device flexibility and durability, especially in maintainingtight tolerances, are essential to attaining planned theoreticalperformance levels.

The considerations listed above are most effective when fitting to heartfeatures of relatively small radius of curvature. For example, it isdifficult to fit regions between the right ventricle and the apex. Agood fit is possible by inserting a pad between the heart and the deviceto fill the gap. However, this adds complexity to the design as well asrequires a rather precise orientation of the device with respect to theheart. Frequently it is the more subtle features which contributesignificantly to device performance. One ideal solution is to form, insitu, a pad between the device and the heart which provides for a customfit. The pad can be a segmented or a single pocket that extends over theentire inner surface of the device. The pocket can be filled with ahydrogel that forms a layer between the wrap and the heart. The designprovides for two entrance ports and an exit port affixed to the pocket.The ports are detachable, and the entrance ports are located at oppositeends of the wrap, the exit port at midway between. A mixture of 10%Hypol and 90% water is mixed in a sealed blender type apparatus. Theresultant mixed solution is introduced to the wrap simultaneouslythrough the entrance ports. The pocket is filled until fluid is detectedat the exit port, and subsequently the exit port is sealed and thepocket gently pressurized so as to completely fill all voids between thewrap and the beating heart. The solution sets in 60 seconds, allowingfor the liquid state of the pocket to be in contact with the beatingheart for approximately 30 seconds. The solution sets in a continuousfashion, reaching a maximum durometer reading in about 10 seconds aftersetting begins. The setting time allows for the pocket to be modified bychanges in the heart shape through at least 10 heart cycles, so that thefinal shape is a fit to the mean heart shape. This provides for theleast loading on the device.

The passive system implementation of the present invention can beunderstood by reference again to FIG. 1A. If the multi-layered balloonelements 304 of the wrap 300 shown in FIG. 1A are filled with acompressible fluid and placed around a patient's heart, the device canboth restrain cardiac hypertrophy and mimic the natural resistance ofthe heart tissue to over-expansion. By choosing an appropriate inflationpressure for the balloon elements 304 and then sealing them, the fluidpressure within the balloons can provide a resistance analogous to theFrank-Starling effect exhibited by cardiac tissue.

In such passive systems, if the heart continues to dilate, the balloons304 will flatten more to accommodate the enlargement but the pressureapplied to the heart by the balloons will be greater. Moreover, unlikemesh-type passive girdles, which rely upon an open structure toaccommodate the expansion and contraction of the heart, the multi-layerinflatable structures of present invention permit the use of solid wrapdevices, which are less likely to loss their effectiveness over time dueto tissue in-growth.

In addition, the passive devices of the present invention can beadjusted. For example, if the heart shrinks, the balloon elements can beperiodically filled to a greater extent in order to tighten the wrap andthe pressure applied by the partially inflated balloons will be less.

All reference materials cited herein are incorporated by reference.

What is claimed is:
 1. An implantable control system for activating acardiac pumping unit comprising: an electronic controller forsynchronized release of actuating fluid to a pumping unit and forsubsequent synchronized evacuation of fluid from the pumping unit, aplenum for storage of a non-pressurized volume of fluid of sufficientsize to provide a flow at nearly constant pressure during the releaseinterval, and an energy converter for periodically pumping fluid fromthe storage plenum to the pumping unit and thus attaining a pressurizedstate in the pumping unit, said energy converter pumping toward thepumping unit during systole and pumping away from the pumping unitduring diastole.
 2. The control system of claim 1 wherein the systemfurther comprises an electrode for sensing a cardiac electrical signal.3. The control system of claim 1 wherein the system further comprises: aconduit for delivering and distributing a pressurized fluid to thepumping unit, an implantable battery and a transcutaneous energyreceiver for recharging the battery.
 4. The system of claim 1, whereinthe pumping unit further comprises a collapsible, non-distensible,biocompatible sheath.
 5. The system of claim 4, wherein said pumpingunit comprises: a plurality of elongate inflatable elements each havinga longitudinal axis; and a sheath defining a set of unit cells, eachcell encompassing at least two juxtaposed elongate inflatable members,wherein said sheath nondistensibly secures said elongate inflatableelements in a substantially parallel position relative to each other. 6.The system of claim 5, wherein said sheath has a length extendingbetween a first and a second end and a width substantially parallel withsaid longitudinal axis, said sheath having inner and outer surfacesjoined periodically along said length at connection points disposedbetween said inner and outer surfaces to form a contiguous plurality ofunit cells.
 7. The system of claim 5, wherein said sheath and saidinflatable elements are each constructed of thin, collapsible,non-distensible, and biocompatible material.
 8. The system of claim 5,wherein said plurality of inflatable elements and said sheath areintegrally formed in a single unit.
 9. The system of claim 5, whereinsaid inflation elements have a top end and a bottom end, and wherein oneor more of said top and bottom ends is tapered.
 10. The system of claim5, wherein said first end and said second end of said sheath areconnected to form a paraboloid of revolution.
 11. The system of claim 5,wherein said plurality of elongate elements are filled with a fillermedium to a first amount and wherein the contractile force applied bythe pumping unit is proportional to said first amount of a filler mediumused to fill said plurality of elongate elements.
 12. The system ofclaim 5, further comprising: at least one pocket interposed between theinner surface of the sheath and the heart, the pocket adapted to befilled with a second filler medium at time of implantation to conformsaid pumping unit to the heart.
 13. The system of claim 5, wherein saidsecond filler medium comprises: a flexible, deformable substance thatsubstantially maintains its volume when compressed.
 14. The system ofclaim 1, further comprising: attachment device configured to preventsaid pumping unit from unintentionally adjusting during operation. 15.The control system of claim 1 wherein the system further comprises; aninternal electronic controller programmed to generate activating signalsto control the release of actuating fluid to the pumping unit and forsubsequent synchronized evacuation of fluid from the pumping unit, andan external controller in radio communication with the internalcontroller.
 16. The implantable cardiac assist system of claim 15wherein the system further comprises: an actuator for converting saidactuating signal into a periodic movement of at least one valve member.17. The implantable cardiac assist system of claim 15 wherein the systemfurther comprises: a volume displacement chamber containing a energyconverter and fluid plenums.
 18. The implantable cardiac assist systemof claim 15 wherein the internal controller further comprises a datatransceiver for receiving a data signal from the external controller.19. The implantable cardiac assist system of claim 15 wherein theinternal controller further comprises a data transceiver fortransmitting a data signal to the external controller.
 20. Theimplantable cardiac assist system of claim 15 wherein the system furthercomprises a sensor for sensing electrical signals capable of indicatingreliably the cardiac cycle.